A typical emission scan using a positron emission tomography (PET) scanner starts with the injection of a solution including a tracer into the subject to be scanned. The subject may be human or animal. The tracer is a pharmaceutical compound including a radioisotope with a relatively short half-life. The tracer has been adapted such that it is attracted to sites within the subject where specific biological or biochemical processes occur. The tracer moves to and is typically taken up in one or more organs of the subject in which these biological and biochemical processes occur.
When the radioisotope decays, it emits a positron, which travels a short distance before annihilating with an electron. The short distance, also called the positron range, is of the order of 1 mm in common subjects. The annihilation produces two high energy photons propagating in substantially opposite directions. The PET scanner includes a photon detector array arranged around a scanning area, usually in a ring-shaped pattern, in which the subject or at least the part of interest of the subject is arranged.
When the detector array detects two photons within a short timing window, a so-called ‘coincidence’ is recorded. The line connecting the two detectors that received the photons is called the line of response (LOR). The reconstruction of the image is based on the premise that the decayed radioisotope is located somewhere on the LOR. It should be noted that in fact the annihilation occurs on the LOR and the decayed radioisotope is a positron range removed from the point of annihilation. The relatively short positron range may be neglected or can be compensated for in the reconstruction. Each coincidence may be recorded in a list by three entries: two entries representing the two detectors, one entry representing the time of detection. The coincidences in the list can be grouped in one or more sinograms. A sinogram is typically processed using image reconstruction algorithms to obtain volumetric medical images of the subject.
The detector array of a typical PET scanner does not rotate during an acquisition and is generally arranged in two or more banks of stationary detector rings. Alternatively, the detectors may be arranged in a non-ring-shaped pattern. In most detector configurations there will be directions in which coincidences are not detectable due to the geometry of the detector array, since the scanner has a finite field of view and there may be blind spots due to gaps between the detectors.
To obtain quantitative results from a PET scan, the processing must, among others, take into account the attenuation of the photons within the subject. An estimate of the attenuation may be obtained by making test scans in which one or more positron emitter rod sources are arranged in the scanner. The sources may be made of a material such as 68Ge, which emits dual annihilation photons. Conventionally, two scans are used to derive the attenuation estimate, a blank scan in which the subject being scanned is not present in the scanning area and, typically, the scanner is empty except for the presence of the sources, and a transmission scan in which both the subject and the sources are present in the scanning area. The results of the blank scan are then divided by the results of the transmission scan, providing an attenuation sinogram. The attenuation sinogram can be used to correct the emission scan of the subject for the effects of attenuation.
Another image degrading factor is the scattering of annihilation photons within the subject. Compton scattering is the dominant mechanism of interaction of the photon in the human tissue. The ratio of scattered photons to the total number of photons detected may be up to 50%. The two photons of an annihilation do not in general travel in opposite directions any more after scattering. Hence, the decayed radioisotope will in general not be located on the LOR of a scatter event. Scatter events therefore degrade the image and the detection data are preferably corrected for scatter. A coincidence without scattering resulting from a single annihilation, where the radioisotope lies on the LOR, is called an unscattered event.
A method for scatter correction is known from the article ‘New, Faster, Image-Based Scatter Correction for 3D PET’ by C. C. Watson in IEEE Transactions on Nuclear Science, Vol. 47, No. 4 (2000) pages 1587-1594. The method reconstructs an image from the detection data that are not corrected for scatter. An estimate of the scatter is calculated using an estimate of the image and of the attenuation. The method calculates a probability for a single-scatter event by tracing two photons resulting from an annihilation, one of which is scattered between the point of annihilation and the detector. The attenuations of the two photons along their paths to the detectors are multiplied with the detector cross sections and the scatter cross-section. A single-scatter estimate is obtained by integrating the probability over all annihilation points and scatter points. A scatter image is reconstructed from the scatter estimate and subtracted from the image derived from the detection data. The resulting image is scatter corrected. An estimate of first and higher order scatter is determined by applying the method for calculating the single-scatter estimate to the non-scatter-corrected image.
A disadvantage of the known method for scatter correction is the inaccuracy of the estimate for the higher-order scatter. It is an object of the invention to provide a method for calculating an improved estimate of higher-order scatter, computer software, a data carrier and a scanner system for carrying out the method.